Ultrafast MRI system and method

ABSTRACT

A Magnetic Resonance Imaging (MM) system, called ULTRA (Unlimited Trains of Radio Acquisitions), can operate with essentially no magnetic gradient reversals. Each of a multitude of small receiver coils arranged in a 3D array around the imaging volume simultaneously acquires MR signal from the entire volume. This greatly increases the rate of MR signal acquisition and allows a full MR image to be reconstructed in as little as 1 millisecond. Both electrical and audible noise is greatly reduced.

REFERENCE TO RELATED APPLICATIONS

This patent application is a continuation-in-part of parent applicationSer. No. 14/671,528, which is scheduled to issue Oct. 18, 2023 as patentSer. No. 10/107,833, and incorporates by reference said parentapplication.

FIELD

This patent specification is in the field of imaging, such as medicalimaging, and more specifically relates to rapid three-dimensional (3D)magnetic resonance imaging (MRI).

BACKGROUND

Several references are identified by numerals in parenthesis in thispatent specification and are hereby incorporated by reference. The fullcitations are listed at the end of the specification. References [18]and [19] are papers by the inventor and persons working under hisdirection related to the subject matter of the paper.

Real-time MR imaging could exert a profound influence on neuroscience inthe future by enabling the direct visualization of neuronalinteractions. For now, however, the known practical embodiments of MRIrequire at least some degree of gradient encoding, and this in turn setsa lower limit of about 100 ms for volume acquisition.

In the original formulation of MRI by Lauterbur [1], spatial encoding isachieved by applying successive magnetic field gradients to the imagingvolume. Each new gradient is associated with a different radiofrequency(RF) excitation, and RF-induced echoes form a line in k-space,discretized into N elements, where N is the dimension of the imagematrix. After N progressively increasing gradients, and N echoes, thereare N lines in k-space, and the N×N k-space matrix is subjected to a 2-DFourier transform, rendering the N×N image matrix. In echoplanar imaging(EPI) Mansfield [2] showed that spatial encoding could be achieved bytrains of gradient reversals after a single RF excitation. An advantageof using gradient reversals is speed; a disadvantage is sensitivity tosusceptibility changes and magnetic field inhomogeneities.

More recently studied parallel MRI (pMRI) uses the spatial sensitivitiesof multiple receiver coils (detectors) arranged around the object forspatial encoding of voxels in a single slice, thus reducing the need forgradient reversals and RF pulses [3, 4]. pMRI images are discussed byKelton, Magin and Wright [5], and by Ra and Rim [6]. pMRI modifies theLauterbur-Mansfield approach to include multiple receiver coils, so thatif the number of coils is n, then the number of gradient reversals is afactor n times smaller. According to such studies of pMRI, many of theconventional 180° RF pulses can be replaced by short trains of gradientreversals, with an acceptable change in image quality. Thus, pMRI canembrace some of the advantages of EPI, without some of thedisadvantages. Because gradient switching is generally several timesfaster than RF excitations, image quality can be adequately maintainedwhile speed is increased somewhat.

The pMRI initially proposed by Hutchinson [3,4] is for single shotimaging that ignores sources of noise in the surface coils. Subsequentanalyses of detector arrays by Roemer [7] and Ocali [8] account to someextent for the effects of noise, with algorithms for conventionalgradient encoding. A concept of “ultimate signal-to-noise,” [8] can beconsidered, by which was meant that, at least for gradient encoding,large numbers of small receiver coils—when properly configured—can bemore efficient than a single receiver. Sodickson [9], with SMASH, andPruessmann [10], with SENSE, proposed merging gradient-encoding andspatial sensitivity encoding. These merging techniques can be designated“hybrid” since they use both methods of encoding.

Merging of the two fundamental ways of encoding comes at a considerableprice, because for the hybrid techniques ultimate signal-to-noise isreached not with large numbers of coils, but with small numbers(typically 4). Signal-to-noise (SNR) in hybrid pMRI is reduced by 3sources: (i) electrical noise, (ii) reduced numbers of echoes, and (iii)a critical geometric factor, g, as discussed by Ohliger [11] andWiesiger [12]. Sources of electrical noise include preamplifiers,coupling between adjacent coils, and body thermal noise. In addition,there is noise from the eddy currents in the body surface due not onlyto the gradient reversals but also to re-radiation of signal by thecoil, which is both a receiver and a dipole radiator. These sources ofnoise can be reduced, however, and for conventional gradient encodingmultiple small coils have been shown to be more efficient than one largecoil [7,8] so there is no insurmountable noise problem inherent to smallcoils. When multiple coils are used to encode, however, there can be anadditional problem due to the inefficiency of spatial sensitivityencoding, compared with gradient encoding, when the coils are large.This in turn is due to the slow spatial variations of the spatialsensitivity of each coil. In addition, since there are fewerphase-encoding steps SNR is further reduced. The signal-to-noise ratiois now given by [11,12]:SNR=SNR_(Full)/(√R)g  (1)where SNR_(Full) is the SNR for gradient-only encoding, R is theacceleration factor (in this case the number of detectors) and g is thegeometric factor. The loss factor √R represents the reduced SNR due tothe reduced number of phase encoding steps and could not be mitigatedabsent phase encoding. The g-factor is intrinsic to the geometry of thecoil array and is a measure of the capacity of the array to compensatefor reductions in gradient encoding. For very small numbers of detectorsg is close to 1, but as the number of coils increases, and the number ofphase-encoding steps correspondingly decreases, g suddenly becomes largeand the images are degraded.

It has generally been accepted that for practical purposes this singlelimitation means that for hybrid techniques acceleration factors need tobe at most about 4, so that the rate of acquisition of SNR is aboutdouble that of non-parallel techniques. The reason g>1 with increasingaccelerations is that the transformations used to obtain the image inconventional pMRI are non-unitary, and the reason for this is that thesolutions to Maxwell's equations are “smooth,” by which is meant thatfor large detectors the spatial dependence of the radio field is not asgranular as the spatial dependence of the gradients used forconventional encoding. This in turn means that large groups of adjacentpixels may have similar spatial sensitivity profiles.

As R increases, and the number of gradient-encoding stepscorrespondingly decreases, more and more of the burden of encoding isshouldered by the receiver array. For R=N n=4, only 25% of the encodingis gradient-based. With more detectors there is an unsupportablereliance on spatial sensitivity encoding and this leads to rapid imagedegradation when R>4. The spatial sensitivity maps can be modeled forillustrative purposes by taking, as an approximation to the solutions ofMaxwell's equations, only the 1/r³ dependence of the near field dipole.This view of an inherent limitation of all constructions of parallel MRIhas been challenged by Keil and Wald [13], with a theoretical evaluationsuggesting that the spatial sensitivity of small coils could contributemore in spatial encoding in hybrid pMRI while still using some gradientencoding. The original concept for single shot, single slice imagingfree of gradient reversals [3,4] was tested several years ago byMcDougall and Wright [14] using a 64-channel coil with an accelerationfactor of 64. However, this was not for volume encoding but only for asingle slice. Proposals have been published to employ static magneticfield gradients produced by thin magnetic films to encode flow [15], touse two or three RF phase gradients in an arrangement free of magneticgradients [16], and to use a flexible array coils populated withmultiple coils in multi-slice-multi-echo sequences that inherently relyon magnetic field gradients [17].

The prior commercial embodiments of pMRI known to the inventor hereinrequire a plurality of magnetic gradients and/or gradient reversals.This can generate a large amount of noise, first RF and other electricalnoise, and second audible noise. This means that image signal-to-noiseratio is lowered, and that the machine can be extremely loud (up to 120dB).

This patent specification recognizes and addresses these and otheraspects of prior MRI work by providing a new, radically differentapproach.

SUMMARY OF THE DISCLOSURE

This patent specification describes a new approach using an MRI systemto do volume imaging that does not rely for 3D spatial encoding onmagnetic gradient switching or additional RF pulses and departs inimportant ways from the known proposals for hybrid pMRI.

One example of the new approach comprises a gantry with a magnetconfigured to generate a main magnetic field B₀ in an imaging volume, agradient field generator configured to generate a steady gradient fieldg in the imaging volume, a radio-frequency (RF) pulse generatorconfigured to apply an excitation RF pulse to the imaging volume, and amultitude of small MR signal receiving coils arranged in athree-dimensional (3D) array surrounding the imaging volume andextending along the B₀ field. Each of the receiver coils is configuredto simultaneously receive RF energy from the entire imaging volumeduring MR signal acquisition and output respective MR signals. An MRsignal acquisition facility acquires the MR signals and acomputer-implemented processor applies image reconstruction algorithmsto the MR signals and thereby generates a 3D image of an object in theimaging volume, displayed as such or as two-dimensional (2D) imagesderived therefrom. The MR signal from each of the coils comprises a timesequence of overall peaks and valleys that can extend over a period ofthe order of seconds and even minutes in response to the RF pulse,without requiring additional MR signal-encoding RF pulses or MRsignal-encoding gradient reversals for spatial encoding or rephasing.

The 3D array of receiving coils comprises M coils along the direction ofthe B₀ field and N coils transverse to B₀. Spatial encoding is based onspatial sensitivity parameters of the receiving coils. The gradientfield generator is configured to generate the steady gradient field g asa single gradient field maintained throughout the MR signal acquisition,and the acquisition facility is configured to acquire the MR signals inthe substantial absence of gradient field reversals or additional RFpulses, although inhomogeneities in the B₀ field can be compensated witha few gradient reversals or additional RF pulses. The imagereconstruction can involve applying a one-dimensional Fourier Transformor Fourier Series process to the MR signal from the respective coils togenerate transformed MR signals and applying a matrix multiplicationprocess to the transformed MR signals using a matrix related to spatialsensitivities of the receiving coils.

The new approach also comprises carrying out an MRI process comprisingapplying a main magnetic field B₀, a steady gradient magnetic field g,and an excitation radio-frequency (RF) pulse to a three-dimensional (3D)imaging volume, acquiring RF energy for the entire imaging volume witheach of a multitude of receiving coils arranged in a 3D array around theimaging volume during MR signal acquisition in the substantial absenceof gradient reversals or subsequent RF pulses, applying imagereconstruction algorithms to the MR signals to thereby generate a 3Dimage of an object in the imaging volume, and displaying the 3D image assuch or as two-dimensional (2D) images derived therefrom.

An embodiment of another aspect of the new approach is a computerprogram product embodied in a non-transitory form in a computer-readablemedium and comprising instructions that, when executed by a computersystem, cause the system to carry out the steps of acquiring magneticresonance (MR) signals for an entire imaging volume generated from eachof a multitude of receiving coils arranged in a 3D array around animaging volume during MR signal acquisition in the substantial absenceof gradient reversals or RF pulses subsequent to an initial RF pulse,applying image reconstruction algorithms to the MR signals to therebygenerate a 3D image of an object in the imaging volume, and displayingthe 3D image as such or as two-dimensional (2D) images derivedtherefrom.

Because each of the receiving coils receives MR signals from the entire3D imaging volume and because of the substantial absence of gradientfield pulses, the output of each of the coils is remarkably superior tothe signals that a single receiving coils outputs in a conventional MRIsystem, or each of the few coils in a known pMRI system outputs. Forexample, the MR signals from each of the small coils is comparable inquality to the signals received by a single coil is in a conventionalMRI system. The implications of this include that for an image ofquality comparable to that from a conventional MRI system, the newsystem described in this patent specification can (i) acquire MRIsignals over a much shorter period, (ii) use much lower strength mainmagnetic field, (iii) greatly improve the SNR (signal-to-noise ratio) ofthe images, and/or (iv) create much higher resolution images forcomparable or even lower main magnetic field and/or for shorter MRsignals acquisition times. For example, the SNR of MRI signals from thenew system can be at least an order of magnitude better that than for MRsignals obtained under otherwise similar conditions from a currentlycommercially available 1.5 T MRI system.

The benefits of needing much lower main magnetic field includeconveniently making the new MRI systems essentially open-MRI, forexample by making the magnet opening very large or making the magnet inportions that are spaced apart by several feet or even meters. As anextreme case, the new system can eliminate magnets as a source of themain magnetic field and rely solely on the Earth's magnetic field as asource of the main magnetic field B₀. The Earth's magnetic field can beroughly 10⁻⁴ times the strength of the main magnetic field in aconventional MRI scanner, but the stronger MRI signals of the multitudeof small coils in the new system can largely compensate for this,assuming appropriate measures are taken to ensure adequate uniformity ofthe Earth's magnetic field acting on the imaging volume.

In addition, because the new system does not rely on gradient pulses ora gradient pulse sequence for MR signals acquisition, and becausegradient pulses are a highly significant source of electronic noise thatcauses image degradation, the MR images from the new system can be muchless noisy compared to a conventional MRI system. This is so becausemuch, if not most of the electrical noise in the MR signals in aconventional MRI system is related to gradient switching. Elimination ofgradient switching, or near elimination, can reduce the electrical noiseto zero or near zero. As a result, the SNR of the new MRI system can beorders of magnitude better than in a conventional MRI system. Inaddition, complete or substantial elimination of gradient pulses canreduce or eliminate the loud acoustical noise that many patients findtroublesome.

As an example, if the 3D array of small coils is n×n coils (n rings eachof n receiver coils), and each coil is A/n of the area A of the largerreceiver coil of a conventional MRI system, the MR signals received bythe small coil would have a strength much greater that 1/n² of thestrength of the MR signals received the larger coil. (“n” is the numberof small coils.) This is so because the smaller coil in effect “sees”all the MR signals from all of the 3D imaged volume, all the time. Incontrast, in conventional MRI systems, each echo encodes for one line ink-space and the signal is correspondingly reduced by a factor of n².

BRIEF DESCRIPTION OF THE DRAWING

FIG. 1a illustrates in block diagram form a portion of an MRI systemaccording to one embodiment described in this patent specification; FIG.1b illustrates a relevant coordinate system; FIG. 1c illustrates aportion of an array of receiver coils and illustrates in block-diagramform the RF pulse generator, the source of the steady magnetic gradient,an MR signals acquisition unit, the image reconstruction processor, andthe image display unit of a complete MRI system; FIG. 1d illustrates inperspective view a portion of the array of the 3D receiver coils; andFIG. 1e illustrates a stack of 2D arrays of receiver coils.

FIG. 2 illustrates an array of receiver coils when unfolded andflattened.

FIG. 3 illustrates an MR signal generated by a receiver coil.

FIG. 4 illustrates a prior art arrangement of receiver coils.

FIG. 5a illustrates an essentially or virtually open-MRI configurationwith a very large diameter magnet bore, and FIG. 5b illustrates anessentially or virtually open MRI configuration with a main magneticfield generated by widely spaced apart magnetic components that can befree-standing or integrated in room walls or floor and ceiling.

DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS

An example of a preferred embodiment can operate by using some of thebasic components of a conventional MRI system, such as a magnetgenerating a main magnetic field B₀ in an imaging volume, a source of RFexcitation pulses, a source of an a magnetic gradient, a computerconfigured to process MR signals and to control the overall operation ofthe system, and a workstation to format and process MR images and allowoperator control. The new approach described in this patentspecification adds other components, and programs the computer in thesystem to carry out different operations. The new components include adrastically different 3D array of very small receiver coils surroundingthe imaging volume where the output of each of the small coils is an MRIsignal comprising contributions from the entire imaging volume. The newapproach further includes modifications in the control over, and ifdesired the structure of, the RF pulse and gradient generators, andproviding a different facility for computer-processing the output of themultitude of small receiver coils.

One example of a preferred embodiment is illustrated in block diagramform in FIGS. 1a-1e and comprises an MRI gantry 10 that can be the sameas or similar to MRI systems currently offered commercially by companiessuch as Siemens Healthcare and GE Healthcare. For example, gantry 10comprises a magnet 102 generating a main magnetic field B₀, using asuperconducting magnet surrounding a cylindrical imaging volume 104,where the main field B₀ extends in a z-direction of a Cartesian xyzreference in which the z-direction is along the length of a patienttable 110 and the x,y directions are perpendicular to the table length,RF pulse generator 106, steady gradient field generator 108 generating asteady magnetic gradient g, an MR signal acquisition facility 115, andcontrols over the components and their operation.

Unlike conventional MRI systems, gantry 10 includes a 3D array 112 of amultitude of very small receiver coils (detectors) 114, which arrayextends both in the direction of the B₀ field and in a transversedirection to thereby surround the imaging volume. Also unlikeconventional MRI systems, gantry 10 is configured to generate a steadygradient magnetic field g extending, for example, in the x-direction (ag_(x) field) but does not require gradient field reversals for spatialencoding as in a conventional MRI system although optionally a fewgradients or reversals could be used for a different purpose, namely, tocompensate for inhomogeneities in the B₀ field. Still in addition,gantry 10 differs from conventional MRI systems in that its RF pulsegenerator is configured to generate a single RF excitation pulse, suchas a 90° excitation pulse, and need not generate additional RF pulsesduring MR signal acquisition for purposes such as spin reversals, exceptoptionally may provide several RF pulses for a different purpose,namely, to compensate for inhomogeneities of the B₀ field. Forconciseness, this patent specification does not provide details of theconventional MRI components and processes, so as not to obscure the newcomponents and processes that are described in greater detail below.

The 3D array 112 of receiver coils comprises a multitude of very small,loop-shaped receiver coils 114, each for example of the order of severalmm along each side. 3D array 112 can be generally cylindrical in shape,surrounding the imaging volume, although it can deviate somewhat from acircular cross-section, for example to an oval shape or a morerectangular shape with rounded corners. The receiver coils 114 arearranged in generally round 2D arrays 116, which 2D arrays can bestacked next to each other in the B₀ direction (the x-direction). The 2Darrays 116 surround the imaging volume. A 3D array 112 can be formed,for example, by a 3D printing technique through that embeds theconductive elements of each coil (and related conductive lines to thecoils) in a material that is not electrically conductive, such as apolymer, and is tubular or is initially flat but later is rolled into atubular shape to surround the imaging volume of the MRI system. Thenumber of such 2D arrays 116 is related to the desired length of the 3Dimaging volume 104 along the B₀ direction and the desired spatialresolution in that direction. For example, if the 3D imaging volume 104is considered as a stack of slices each 1 mm thick in the B₀ direction,and if the imaging volume is 20 cm long in the B₀ direction, then therecan be a stack of approximately 200 2D arrays 116 of receiver coils 114forming the 3D array 112. However, there need not be a one-to-onecorrespondence between the number of 2D arrays of receiver coils and thenumber of actual or virtual slices of the imaging volume.

FIG. 1c illustrates a spatial arrangement of receiver coils 114 in amagnified partial view, where one of the coils is labeled 114 mn toillustrate an example of “m” 2D arrays 116 of coils 114 and “n” coils114 in each 2D array 116. The 2D arrays may all have the same number ofcoils per array but in another embodiment, some 2D arrays may havedifferent numbers of coils. FIG. 1c also schematically illustrates thatthe MR signals from coils 114 go to a MR signals acquisition facility orunit 120 and are processed into MR images in image reconstructionprocessor 122 for display on display 124 (or for other purposes such ascomputer image analysis or manipulation, storage, etc.). FIG. 1c alsoillustrates RF generator 106 that produces the required RF pulse orpulses and a gradient generator 108 that maintains a steady magneticgradient during MR signals acquisition.

FIG. 2 illustrates an arrangement of small receiver coils 114 when the3D array 112 of coils is unfolded and flattened for easiervisualization. There can be M 2D arrays 116 of coils 114 in the B₀direction, which can but need not correspond to M actual or virtualslices of the imaging volume. Each 2D array 116 can comprise N coils114, where M×N can be 256×256 in one non-limiting example. M and N arepositive integers much greater than unity that may but need not be thesame. Each of the coils 114 is identified by the reference numeral 114mn, to indicate that it is the m-th coil in the B₀ direction in the 3Darray 112, and the n-th coil within the 2D array 116 of coils 114 in adirection transverse to the B₀ direction. Each detector 114 is depictedas a single loop. As an example, for a spatial resolution of 1 mm and afield of view of 250 mm, each of the loops 114 can be approximately 3 mmperpendicular to B₀. Preferably each coil 114 is less than 1 cm in eachdimension; more preferably, each is less than 5 mm in each dimension;still more preferably each coil is 3 mm or less in each dimension.

Gantry 10 excites an object in the imaging volume 104 with a single RFpulse, for example a 90° RF pulse, while the object is in the imagingvolume and an initial steady gradient field g is being applied to theimaging volume. In the example of a 1.5 T MRI magnet with the B₀ fieldin the z-direction, the initial gradient field can be in the x-directionand can have a typical strength of 45 mT/meter. Unlike in conventionalMRI, the gradient magnetic field g need not change during MR signalsacquisition for the purpose of generating or acquiring MR signals, andthere need not be other gradient fields and reversals thereof forspatial encoding. In addition, there needs to be only an initial RFexcitation pulse for generating MR signals. The only additional RFpulses and gradients or gradient reversals that can be employed, ifdesired, serve to compensate for distortions of MR signals from thecoils 114 due to inhomogeneities in the main magnetic field or possiblyin the steady gradient field. This compensation can serve to rephasedephasing due to B₀ field inhomogeneities but is not required forexciting the object for the purpose of generating MR signals. Thus, theadditional RF pulses and gradients/reversals, if any, can be calledfield inhomogeneity compensating pulses or gradients, or simplycompensating RF pulses or gradients.

FIG. 3 illustrates an MR signal S that a single receiver coil 20generates in response to the initial excitation RF pulse. Notably, theresponse MR signal S does not simply rise and fall, as a conventional MRecho in a conventional MRI system, but rather oscillates, formingrepeated overall peaks and valleys without further RF excitation of theimaging volume. This oscillation can go on indefinitely if the T1 and T2parameters of the object in the imaging volume did not decay inamplitude with time. In practice, the oscillations can provide useful MRsignal of the order of seconds and minutes consistent with the quantumdecay (spontaneous emission) of the excited state. Even without decayand without gradient switching, the amplitude of the MRI signal S ofeach of the small coils goes to nominal zero every τ=(FOV)/(256γg_(x))seconds, where FOV is the field of view, γ is the gyromagnetic ratio,g_(x) is the magnetic field gradient in the x direction, and 256 is theimage matrix dimension in pixels

The new process is explained below in greater detail. As noted, it makesuse of a massively parallel, 3D array 112 of very small receiver coils(detectors) 114 arranged around the object being imaged. Unlike theprevious proposal for single-slice pMRT (3, 4), each of the small coils20 in the new approach simultaneously acquires MR signals from theentire 3D imaged volume. This leads not only to markedly reduced imagingtimes, but also to marked increase in MR signal per unit of time.

While in theory current MRI systems could acquire signal from the entirevolume simultaneously by successive application of gradients in the x, yand z directions, and could analyze the signal with a 3D Fouriertransform, in practice after each successive gradient in the z-directionthe entire set of spins would have to be allowed to relax, and thenre-excited for further signal acquisition. Otherwise the T2 dephasingwould quickly eliminate the useful MR signal. Furthermore, the requiredperiods of relaxation would greatly extend actual acquisition times.

In contrast, in the new system described here, which can be identifiedby the label ULTRA, there is a steady, single gradient g employedthroughout the MR signal acquisition for MR signals acquisitionpurposes. This single magnetic gradient g need not be switched or evenvaried, and in one preferred embodiment is never switched or variedduring MR signal acquisition. As a result, the technique is essentiallynoise-free both electrically and acoustically. In practice, to overcomeor reduce the effect of inhomogeneities of B₀, the ULTRA processpreferably uses a periodic imposition of inhomogeneity-compensating 180°RF pulses, for example approximately every 10 ms. The resultingelectrical noise can be a low-frequency hum at approximately 100 Hz thatcan be filtered out, if desired, for example with a suitable low-passfilter.

The massively parallel 3D array of small coils (detectors) 114, M×N innumber, arranged in a generally cylindrical or round fashion around theobject in the imaging volume, also is very different from the receiverin the earliest description of pMRI [3, 4], where only a single slicewould be excited at any given time in order to provide sufficientamplitude of the detected MR signal. 3D image acquisition and 3D imagereconstruction were not proposed or contemplated. The detector arraythat was believed to be most consistent with maximal signal in theearlier pMRI proposals [3,4] consisted of N long parallel loops on acylinder, arranged in a single layer with long axis parallel to B₀, asillustrated in FIG. 4 (where the plane of the imaged slice isperpendicular to the lengths of the coil loops).

In the new ULTRA system, it has been unexpectedly discovered that the MRsignal per unit time is very significantly increased, if not maximized,by acquiring MR signals from the entire imaged volume simultaneously ateach of the small receiver coils 114, provided the receiver coils(detectors) are made much smaller than in conventional MRI systems. Theuse of such much smaller receiver coils 114 according to the newapproach paradoxically enables a much higher rate of acquisition of MRsignal and/or strength of the acquired signals.

For simplicity, and to illustrate the new system and method, firstconsider a hypothetical case in which the parameters T1 and T2 of anobject in the imaging volume 104 do not decay with time, so that thereis no decay of MR signal with time and the image is defined only by thespin density ρ_(ijk), where (i,j,k) are the quantized (pixelated)coordinates corresponding to (x,y,z) voxels in an object in the imagingvolume. The principal steps in an example of the new process can be:

-   1. Apply a single steady gradient g in a fixed direction, say g_(x),    which remains fixed throughout the MR signal acquisition;-   2. Excite all spins in the entire volume of the object being imaged,    simultaneously, with a 90-degree RF pulse.    The MR signal S_(mn;ijk)(t) in a small coil (detector) 114 _(mn) due    to volume element (i,j,k) at time (t) in the object in the imaging    volume is then    S _(mn:ijk)(t)=R _(mn;ijk)ρ_(ijk) e ^(iωt)  (2)    where ω=xg_(x), x is distance in the x-direction, and R is a matrix    representing detector spatial sensitivity of coils 114 in 3D array    112, based on a near-field dipole spatial dependence of 1/r³, where    r is the distance between a given spin and a given small coil    (detector) 114, and an angular dependence as given by Maxwell's    equations. Such a spatial dependence can be modified at high field    strengths and low wavelengths, but this is known in MRI technology    and the way to do it is within ordinary skill in that technology.    Since ω is proportional to x, it can simply replace x and Eq. (2)    can be rewritten:    S ^((ω)) _(mn;jk)(t)=R ^((ω)) _(mn:ijk)ρ^((ω)) e ^(iωt)  (3)    where the left hand side is the amplitude of the MR signal in small    coil (detector) 114 _(m,n) due to the volume element at location    (j,k) in the yz plane defined by frequency ω (or, in other words,    the yz plane defined by fixed x). Then:    S ^((ω)) _(mn)(t)=Σ^((ω)) _(mn;jk)ρ^((ω)) _(jk) e ^(iωt)  (4)    -   j,k        Let the one dimensional Fourier transform of S(t) be        ^((ω)), where        ^((ω)) is a matrix representing the ω^(th) component of the set        of Fourier transformed MR signals in all small coils 114 (all        detectors).        With this definition, Eq. (4) can be rewritten as:        ρ^((ω)) ={R ^((ω))}⁻¹        ^((ω))  (5)        which in words means the following: the two-dimensional matrix        representing the spin densities within the plane defined by        frequency ω, is the product of the inverse spatial sensitivity        matrix of the small coils 114 in 3D array 112 and the matrix        representing the Fourier transform of the time dependent signals        in all small coils (detectors) 114.

Eq. (5) above has some similarities to Eq. (15) of reference [4] butexplains a process that is not described or suggested in reference [4].The left-hand side of Eq. (5) above represents a two-dimensional imageplane in a process based on a 3-dimensional structure where all imageplanes are acquired simultaneously. In contrast, the left-hand side ofEq. (15) of reference [4] represents a line of image data in the contextof a 2-dimensional structure where all lines in only a single imageplane are acquired simultaneously. Thus, Eq. (5) above achieves theresults described in this patent specification in the context of theradically different structure of a 3D array 112 of receiver coils 114each of which is made much smaller than in reference [4], and the entireimaging volume can now be imaged with one excitation.

Eq. (5) above embodies a key concept of the new approach anddemonstrates that the entire imaging volume, including all planes ofequal frequency defined by a single linear and steady gradient g, can bedecoded uniquely by first performing a Fourier transform of the MRsignal in each small detector 114, and then multiplying the transformedsignals by an inverse of the matrix of spatial sensitivities of thereceiver coils 114. It can now be seen that each small detector loop 114receives MR signal information from the entire imaging volume.Therefore, all the MR signals, from all the spins, all of the time, aredetected by each receiver coil 114 in the entire set of M×N detectors114. This vastly increases the rate of signal acquisition and puts ULTRAin a unique category, since in known conventional embodiments of MRI atany given time only a partial volume of spins gives rise to MR signals.It therefore follows that ULTRA is not only faster than known previousproposals and embodiments but also gives rise to much greater rates ofMR signal acquisition. Put another way, if the entire set of M×Ndetectors 114 in ULTRA were joined to form one large conventionaldetector, this would capture a much greater MR signal than a largeconventional detector employing gradient reversals, since all spins inthe object volume would be contributing to the signal at all times.Nevertheless, such an arrangement would not make instant 3-dimensionalimaging possible, because there is only one coil. ULTRA takes advantageof an inherently very high MS signal by distributing it throughout avery large number of small coils, and then reconstructing it. Withreference to FIG. 2, where the matrix of receiver coils 20 is mapped ona two-dimensional plane, it now becomes apparent that ULTRA can beconsidered a holographic technique in some respects. That is, the entirethree-dimensional volume of spin density in the imaging volume can besimultaneously and uniquely encoded in a two-dimensional plane (on whichthe generally cylindrical array of small coils can be unfolded).

The MR signal processing described above can be carried out in aprocessor such as currently used in commercial MRI systems butprogrammed differently to carry out the processing, using a program thata person of ordinary skill in MRI and computer programming can writewithout undue experimentation based on the explanation herein. Such aprogram can become a computer program product when stored in anon-transitory manner in computer-readable media such an optical disc, athumb drive, or a magnetic disc.

Referring again to FIG. 3, as noted above the oscillating MR signalwould go on forever for infinite T1 and T2, or in practice until thequantum decay (spontaneous emission) of the excited state ends, whichcan be on the order of many seconds and even minutes. The signal isperiodic, and even without decay, and even without gradient switching,its amplitude (vertical axis in FIG. 3) would go to nominal zero everyτ=(FOV)/(256γg_(x)) seconds, where FOV is the field of view, γ is thegyromagnetic ratio, g_(x) is the magnetic field gradient in the xdirection, and 256 is the image matrix dimension. Therefore, in thespecial, hypothetical case of infinite T1 and T2, the signal is aperiodic function such that within each period τ the amplitude would goto nominal zero at least once, as dictated by g_(x). In real life, notonly are T1 and T2 finite, but for short time scales the MR signal willdecay because of inhomogeneities in B₀. A rough estimate of the latterdecay time can be made by assuming an inhomogeneity of p parts permillion. In the example of a 1.5 T (Tesla) MRI system, where thefrequency is 65 MHz, and for p=1, the signal would begin to decay in asignificant way after 1/65 seconds, or approximately 15 ms. To correctfor this, assuming longer MR signal acquisition is desired, a periodicinhomogeneity-compensating 180° RF pulse can be applied, say every 10 msor some other suitable period. For p=2, in this example the RF pulsewould be applied approximately every 5 ms, etc. The imposition of suchinhomogeneity-compensating RF pulses would lead to a low-frequencyhumming sound, which would be quite different from the high frequency,high decibel sounds of gradient reversals which characterize current MRItechnology and are disturbing to patients. Note that theseinhomogeneity-compensating RF pulses are not used to encode, only tore-phase any dephasing caused by inhomogeneity in B₀.

While the foregoing description pertains to spin density images, inclinical MRI embodiments images can be of interest that are sensitivealso to the spin-lattice relaxation time, T1, and the spin-spinrelaxation time, T2. Such images also can be obtained using the ULTRAprinciples explained in this patent specification. Note first that fromeach broad peak in FIG. 3 an entire 3-dimensional image can be created.If T1 and T2 were infinite, then each broad peak would have the sameheight (MR signal amplitude). As one example, for 10 peaks at a time,comprising say 10 ms of MR signal, the entire signal can be averagedinto one image having high signal-to-noise ratio. The next set ofimages, from 10-20 ms, can likewise be averaged into one image, and soon, until MR signal has been acquired for, say, 10 high-signal imagesets spanning 100 ms. T2 can then be determined for each pixel withinthe imaging volume by fitting it to an exponential decay generated fromall image sets.

Likewise, in this example after 100 ms the spins can be rephased by apartial excitation RF pulse, and the process repeated, giving yet moreMR signal for the T2-weighted images, while at the same time allowingfor an estimation of T1, by the relative amplitudes of the broad peakscompared with the amplitudes after the initial 90° RF excitation pulse.The entire process can be repeated every 100 ms, so that after, say,1000 ms all information pertaining to both T1 and T2 can be obtainedwith high signal-to-noise ratio. Still likewise, similar principles canbe used to create inversion-recovery images or diffusion-weightedimages. For example, inversion-recovery images can be made modifying theconventional 180° RF pulse, followed after a time TI by a 90° RF pulse,so that the 90° pulse occurs, say, at time TI−50τ, with acquisitionbeginning at that time, and continuing until TI+50τ. Diffusion-weightedimages might be obtained in analogy with conventional imaging, byapplying diffusion gradients every τ milliseconds.

FIG. 5a illustrates an MRI system according to the principles describedherein that is essentially or virtually an open-MRI system. Because ofthe high strength and quality of the MR signals in the new approachdescribed above, the main magnetic field can be much lower strength thanin conventions MRI system, and can be provided as one example by amagnet 102 that has a diameter so much greater than of the bore of aconventional MRI system that a patient is much less likely perceive itas a confining, tunnel-like enclosure. For example, the magnet may bespaced from the imaging volume by 3 feet or more in each direction, sothat for a cylindrical magnet the distance from the inner surface of thebore to the imaging volume is 3 feet or more, preferably is 4 feet ormore, and even more preferably is 5 feet or more. While the mainmagnetic field applied to the imaging volume can be much less than in acurrent conventional MRI system, the resulting image can be comparableor better because of the MR signal characteristic in the new systemdescribed above. For further patient comfort and/or to allow observationof the patient during MR signal acquisition, the 3D array of coils 114can be mounted on a transparent or at least translucent structuresurrounding the patient. In other respects, the arrangement of FIG. 5acan be structured and can operate as the ULTRA system described above,with elements 106, 108, 120, 122, and 124, plus any additionalfacilities for further image processing and/or image storage. As oneexample, the main magnetic field can be at least an order of magnitudeless than 1.5 T at the imaging volume, where the imaging volume issufficiently large to encompass a cross-section of the torso of a humanpatient.

Similarly, FIG. 5b illustrates an MRI system that a patient may perceiveas essentially or virtually open-MRI system because the main magneticfield is provided by a magnet in the form of two structures 102 a and102 b that are spaced apart by several feet or meters. For example, thedistance from the imaging volume to the nearest surface of the magnet is3 feet or more, preferably is 4 feet or more, and even more preferablyis 5 feet or more. One or both of structures 102 a and 102 b can bebuilt into a wall or walls of the room, or one of the structures or bothstructures can be free-standing. In all other respects, the arrangementof FIG. 5b is as in FIG. 5a , i.e., it includes elements 106,0108,120,122, and 124 plus any additional facilities for further image processingand/or storage.

In case the Earth's magnetic field is relied on as the source of themain magnetic field, the new system can eliminate the magnet 102altogether and can be as illustrated in FIG. 1d —the entire MRI systemcan be the 3D array 112 of coils 114, RF generator 106 and gradientgenerator 108 acting on the imaging volume, signal acquisition facility120, image processor 122, and image display 124 (plus any desiredadditions such as workstations for further image processing and anyimage storage facilities).

While several embodiments are described, the new subject matterdescribed in this patent specification is not limited to any oneembodiment or combination of embodiments described herein, but insteadencompasses numerous alternatives, modifications, and equivalents. Inaddition, while numerous specific details are set forth in thedescription to provide a thorough understanding, some embodiments can bepracticed without some or all of these details. Moreover, for clarity,certain technical material that is known in the related art has not beendescribed in detail in order to avoid unnecessarily obscuring the newsubject matter described herein. It should be clear that individualfeatures of one or several of the specific embodiments described hereincan be used in combination with features or other described embodiments.Further, like reference numbers and designations in the various drawingsindicate like elements.

The foregoing has been described in some detail for purposes of clarity,but it will be apparent that certain changes and modifications may bemade without departing from the principles thereof. It should be notedthat there are alternative ways of implementing both the processes andapparatuses described herein. For example, the disclosed process isintended to work by using an MR signal acquisition that does not includeany gradient reversals or MR signal-encoding RF pulses after the initialRF excitation pulse. However, a substantial absence of such gradientreversals or subsequent RF pulses may be sufficient, where “substantialabsence” for the purpose of this patent specification includes the useof some gradient reversals and subsequent signal-encoding RF pulses thatstill allow the reconstruction of clinically useful 3D images from MRsignals that are due mainly to the initial RF excitation pulse and aremainly due to the initial gradient that remains steady throughout imageacquisition. Accordingly, the present embodiments are to be consideredas illustrative and not restrictive, and the body of work describedherein is not to be limited to the details given herein, which may bemodified within the scope and equivalents of the appended claims.

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The invention claimed is:
 1. A magnetic resonance imaging (MRI) systemconfigured to simultaneously acquire MR signals from an entirethree-dimensional (3D) imaging volume that is in a steady main magneticfield, using spatial sensitivity parameters of a multitude of smallreceiver coils without a need for spatial encoding by plural gradientreversals or RF pulses in addition to an initial excitation RF pulse,comprising: a gradient field generator configured to apply a steadygradient field g to said imaging volume in said main magnetic field B₀and a radio-frequency (RF) pulse generator configured to apply saidinitial excitation RF pulse to the imaging volume; a multitude of MRsignal receiving coils arranged in a 3D array surrounding the imagingvolume, said array extending along a direction of the B₀ field as wellas transversely to the B₀ field direction; each of the receiver coilsbeing configured to simultaneously receive RF energy from the entireimaging volume during MR signals acquisition and to output respective MRsignals in response to said initial excitation RF pulse and while saidsteady gradient field and main magnetic field are present; an MR signalacquisition facility configured to acquire the MR signals; acomputer-implemented processor configured to apply image reconstructionalgorithms to the MR signals and thereby generate a 3D image of anobject in the imaging volume; and a display facility configured todisplay the 3D image as such or as two-dimensional (2D) images derivedtherefrom.
 2. The MRI system of claim 1 in which the MR signals fromeach of the coils comprise a time sequence of overall peaks and valleysthat extends over a period of the order of at least seconds in responseto the initial RF pulse and without requiring gradient reversals forspatial encoding or RF pulses in addition to said initial RF pulse. 3.The MRI system of claim 1 in which the 3D array of receiving coilscomprises M coils along the direction of said main magnetic field B₀ andN coils in each of a multitude of adjacent planes each transverse tosaid direction.
 4. The MRI system of claim 1 in which the RF pulsegenerator is further configured to generate magnetic fieldinhomogeneity-compensating RF pulses after the initial RF pulse.
 5. TheMM system of claim 1 in which the gradient field generator is configuredto maintain the steady gradient field g throughout the MR signalsacquisition.
 6. The MRI system of claim 1 wherein the acquisitionfacility is configured to acquire the MR signals in the substantialabsence of gradient field reversals.
 7. The MRI system of claim 1wherein the acquisition facility is configured to acquire the MR signalsin the substantial absence of RF excitation pulses, other than magneticfield inhomogeneity correcting pulses, subsequent to the initial RFexcitation pulse.
 8. The MM system of claim 1 in which the processor isconfigured to apply a Fourier Transform or Fourier Series processing tothe MR signals to generate transformed MR signals, and to apply a matrixmultiplication process to the transformed MR signals using a matrixrelated to spatial sensitivities of the receiving coils.
 9. The Millsystem of claim 1 in which the MR signals from each of the coils conformto a succession of overall peaks and at least one of the acquisitionfacility and the processor is configured to combine the MR signals ofsuccessive multi-peak sets of the overall peaks of the MR signals intosuccessive sets of combined signals and the processor is configured toderive an estimate of at least one of T1 and T2 by comparing thesuccessive sets of combined signals.
 10. The Mill system of claim 9 inwhich the RF pulse generator is configured to generate a partialexcitation RF pulse rephrasing spins at a time during the acquisitionand/or combining of MR signals, and the processor is configured toderive an estimate of T1 by comparing combined signals obtained beforeand after the rephrasing RF pulse.
 11. The MM system of claim 1 in whichRF generator is further configured to generate an inversion-recovery RFpulse and the processor is further configured to process the MR signalsinto an inversion-recovery image.
 12. The MM system of claim 1 in whichthe gradient field generator is configured to apply a diffusion-gradientand the processor is configured to process the MR signals into adiffusion-weighted image.
 13. The MRI system of claim 1 furtherincluding a magnet generating said main magnetic field B₀.
 14. The MMsystem of claim 13 in which said magnet is spaced from the imagingvolume by 3 feet or more in each direction.
 15. The MRI system of claim1, in which the main magnetic field is at least an order of magnitudeless than 1.5T and the imaging volume is sufficiently large to encompassa cross-section of the torso of a human patient.
 16. The MM system ofclaim 1, in which the SNR of the MR signals obtained from an object insaid imaging volume is at least an order of magnitude better than thatof MM signals obtained in otherwise comparable conditions by a currentcommercially available 1.5T MM system.
 17. The MM system of claim 1, inwhich the main magnetic field is solely the Earth's magnetic fieldexcept for possible sources of fields correcting for inhomogeneities ofthe Earth's magnetic field at the imaging volume.
 18. A magneticresonance imaging (MRI) method comprising: applying a gradient magneticfield g and an initial excitation radio-frequency (RF) pulse to athree-dimensional (3D) imaging volume in the presence of a steady mainmagnetic field acting on said imaging volume; acquiring MR signals forthe entire imaging volume from each of a multitude of receiving coilsarranged in a 3D array around the imaging volume, during MR signalacquisition in the substantial absence of spatial-encoding gradientreversals or RF pulses after the initial excitation RF pulse; whereineach of said receiving coils receives MR signals generated by the entireimaging volume; applying computer-implemented image reconstructionalgorithms to the MR signals to thereby generate a 3D image of an objectin the imaging volume; and displaying the 3D image as such or astwo-dimensional (2D) images derived therefrom.
 19. The method of claim18 in which the gradient magnetic field g is a single field maintainedthroughout the MR signal acquisition.
 20. The method of claim 18including applying, during MR signal acquisition, at least one ofinhomogeneity-compensating gradient reversals andinhomogeneity-compensating subsequent RF pulses to compensate forinhomogeneities in the B₀ field.
 21. The method of claim 18 includingapplying, during the MR signals acquisition, bothinhomogeneity-compensating gradient reversals andinhomogeneity-compensating subsequent RF pulses to compensate forinhomogeneities in the main magnetic field.
 22. The method of claim 18in which the applying of the image reconstruction algorithm furtherincludes obtaining estimates of at least one of T1 and T2.
 23. Themethod of claim 18 in which the applying of the image reconstructionalgorithm further includes obtaining an inversion-recovery image. 24.The method of claim 18 in which the applying of the image reconstructionalgorithm further includes obtaining a diffusion-weighted image.
 25. Amagnetic resonance imaging (MRI) method comprising: acquiring MR signalsfrom a three-dimensional (3D) object in an imaging field in response toan initial RF excitation of the object but in the absence of subsequentspatial encoding RF excitation pulses and/or spatial encoding gradientmagnetic field reversals; reconstructing a three-dimensional (3D) imageof the object from the MR signals; and displaying the 3D image.
 26. Themethod of claim 25 in which the acquiring comprises spatial encodingusing spatial sensitivities of a multitude of receiving coils.
 27. Themethod of claim 25 in which the acquiring of MR signals comprisesacquiring the signals in the presence of a single magnetic gradient thatremains unchanged throughout signal acquisition.
 28. The method of claim25 in which the signal acquisition comprises acquiring, from each of thecoils, a succession of overall peaks of the MR signal, and wherein thereconstructing comprises reconstructing a 3D image of the object from aset of a single peak from each of the coils.
 29. The method of claim 25including applying at least one of inhomogeneity-compensating gradientreversals and inhomogeneity-compensating RF pulses to compensate forinhomogeneities in a main magnetic field B₀ applied to the object.
 30. Acomputer program product embodied in a non-transitory form in acomputer-readable medium and comprising algorithms that, when executedby a computer system, cause the system to carry out the steps of:acquiring magnetic resonance (MR) signals for an entirethree-dimensional (3D) imaging volume that is in a main magnetic fieldB0, wherein said MR signals are generated from each of a multitude ofreceiving coils arranged in a 3D array around an imaging volume, saidarray extending both along a direction of said main magnetic field B0 aswell as transversely to said direction of said main magnetic field B0,during MR signal acquisition in the substantial absence of spatialencoding gradient reversals or spatial encoding RF pulses after to aninitial encoding RF pulse; wherein each of the receiver coils isconfigured to simultaneously receive RF energy from the entire imagingvolume during said acquiring of said MR signals and to output respectiveMR signals; applying image reconstruction algorithms to the MR signalsto thereby generate a 3D image of an object in the imaging volume; anddisplaying the 3D image as such or as two-dimensional (2D) imagesderived therefrom.